Prior art orthopedic prostheses are used to replace or supplement a portion of the natural bone. Most orthopedic prostheses include at least one surface region intended for secure fixation to the natural bone. Orthopedic prostheses may also include a load bearing surface region that is disposed adjacent to the bone without secure affixation. Many orthopedic prostheses replace a natural joint, such as a knee, hip, shoulder, finger or ankle. Prosthetic joint replacement systems will include two prosthetic components affixed to separate bones and capable of articulating relative to one another. Thus, a component of a prosthetic joint may include an articulating surface region, a load bearing non-articulating surface region and a surface region for direct affixation to adjacent bone.
Prior art prosthetic systems have employed screws, wedge fitting and cement either separately or in various combinations to fix the prosthetic device to the bone. However, cement is known to at least partly deteriorate after an extended time in the body, thereby causing a shifting of the prior art prosthetic device relative to the bone. This movement abrades the cement, the bone and the prosthetic device, resulting in release of wear debris which produces bone death and further loosening. Prosthetic devices relying entirely on screws or wedge fitting also may shift over time in response to changes in the natural bone and/or forces exerted on the prosthesis, and abrasion resulting from such shifting generates wear debris as explained above.
The more recent prior art has included prosthetic devices with surface coatings or modifications that are intended to enhance natural bone ingrowth to the prosthesis for achieving a biological fixation of the prosthesis to the natural bone. More particularly, at least selected surface areas of these prior art prostheses are provided with a macro surface defining pores, fissures, texturing or the like into which the natural bone tissue may grow. This biological fixation is intended to stabilize the prosthetic device relative to the bone and substantially prevent the loosening or shifting which had occurred with the above described earlier fixation means.
The surface treatment to enhance the bone ingrowth to the prior art prosthetic device can be achieved in many ways. A common surface treatment for prior art prosthetic devices is referred to as a porous coating which is applied to the metallic substrate of the prosthesis to define an array of small pores or fissures in the surface of the prosthesis. Other prior art prosthetic systems apply a metallic mesh material to a substrate, such that the mesh material defines the texture into which bone tissue may grow. Still other prior art systems directly modify the substrate to define the surface irregularities, such as small holes, fissures, slots or the like.
Orthopedic prosthetic devices typically are formed from a metallic alloy that will exhibit appropriate strength and flexure in use. Examples of metallic alloys that are currently used for orthopedic prosthetic devices include titanium alloys, such as a titanium aluminum vanadium alloy, and cobalt-chromium alloys, such as a cobalt-chromium molybdenum alloy. Although both titanium and cobalt-chromium alloys exhibit appropriate strength and flexure for most applications, each such alloy has its own unique advantages and deficiencies. For example, cobalt-chromium exhibits desireable hardness, and hence is widely used for prosthetic devices having an articulating surface region. However, cobalt-chromium is very expensive and its desireable hardness characteristics make it difficult to machine. Furthermore, some patients exhibit sensitivity to cobalt-chromium alloys. Additionally, some cobalt-chromium alloys are known to release metallic ions as a result of corrosion after extended exposure to the biological tissue of the human body. These ions are suspected to cause tumors and may have carcinogenic effects.
The amount of ion released from the alloys defining the substrate of prosthetic devices generally has been considered acceptably low. However, it is known that the amount of ion release increases with the surface area of the prosthesis. The relatively recent advent of surface coatings or modifications to promote bone ingrowth results in a very substantial increase in surface area. More particularly, pores, fissures or other such surface irregularities vastly increase the surface area for a given prosthesis as compared to the same prosthesis having a smooth exterior surface. This increase in surface area achieved by the pores, texturing or other surface irregularities has resulted in an increase in the ion release from the alloy from which the prosthesis is made. The ions released from the alloy of the prior art prosthetic device may leach into the body and migrate to areas remote from the prosthetic device. Patients having prostheses with surface treatments for promoting bone ingrowth have been observed to have ions in urine specimens, liver tissue, and other body locations remote from the prosthesis.
In view of these fairly recent findings, some doctors recommend not employing prostheses with macro surface treatments in young patients who may be expected to be exposed to the large surface area of ion releasing alloys for a number of years and who therefor run a greater risk of being adversely impacted by the tumor causing or carcinogenic effects of the ions.
In addition to the preceding negative effects of ion release from prior art prosthesis having surface treatments for promoting bone ingrowth there is a desire to improve the biological fixation between the prosthesis and the natural bone. The ingrowth of bone to the macro surface region of the prior art prosthesis often is not complete. Data suggest that bone will grow into only about 10-20% of many prior art macro surface area. Furthermore, a thin fiber layer may exist between the substantially rigid bone and the pores, fissures or other irregularities of some the prior art macro surfaces. This relatively incomplete bone ingrowth may be due to the above described corrosion reaction and resulting ion release. The effect of the incomplete bone ingrowth may be some movement between the prosthesis and the bone. As noted above, such movement will generate metallic wear debris. The small metallic particles produced by such wear corrode rapidly in view of the relatively large surface area to volume ratio for these particles. As noted above, corrosion results in the undesirable metallic ion release. Thus a more complete biological fixation could reduce the potentially harmful metallic ion release.
Titanium alloy prostheses generally are considered to be much more biologically compatible than the cobalt-chromium alloys. Thus, the sensitivity some patients have to cobalt-chromium prostheses generally is not a problem for titanium alloy prosthetic components. Titanium alloys also are substantially less expensive than cobalt-chromium alloys, but they are not as hard. Thus, titanium alloys prosthetic components are likely to generate wear debris when employed on articulating surface regions of a prosthetic component or on a non-articulating load bearing surface region that is subject to micromovement relative to adjacent bone. The metallic wear debris particles cause further deterioration of both the prosthetic component and the natural bone.
The prior art has included prosthetic components formed from a plurality of different metallic alloys. For example, femoral prosthesis have been provided with a titanium alloy stem and neck and a cobalt-chromium alloy head. The comparatively harder cobalt-chromium head performs well as an articulating surface. The titanium alloy stem and neck achieve better biocompatibility at a lower cost. However, galvanic action is known to be generated at the interface of the titanium alloy neck and the cobalt-chromium alloy head, with a corresponding corrosion and generation of corrosion-related debris. Additionally, load bearing areas of the stem are subject to micromovement relative to the bone, and hence these areas can generate significant wear debris. Furthermore, it would be desireable to provide a less expensive, harder and more biologically compatible head than the cobalt-chromium head in the prior art system.
The prior art also has included attempts to provide coatings on a metallic alloy prosthesis to enhance some aspect of its performance. The coatings have included ceramics which are noted for their hardness and their biological compatibility. These ceramic coatings have been applied to the metallic alloy substrate by known thin film technology. This technology is widely used, for example, in the machine tool arts to enhance the life of cutting tools. Prior art thin film technology typically applies the ceramic coating to a substrate in a vacuum coating chamber. The substrate to be coated functions as the cathode in the chamber, while the anode is formed from the material to be coated onto the substrate. An arc is struck in the chamber, and the substrate to be coated is subjected to high energy ion bombardment from the anode. A gas is then introduced into the chamber. The gas reacts with the ions of the anode and produces an ionic deposition of a highly-adherent ceramic coating onto the substrate. This thin film technology typically is employed to provide an tonically bonded coating approximately 2-4 microns thick. Thicker ceramic coatings have been considered too costly and problematic for application by thin film technology, and hence thin film technology has not been employed for thicker coatings in the machine tool art and have not been carried over into the orthopedic prosthetic art. The problems of using thin film technology to produce a thick coating have been cracking and eventual delamination. It has been believed that cracking occurs in part due to the different stiffnesses of the metallic alloy substrate and the ceramic coating. Thus, ceramic coatings applied by thin film technology on prosthetic devices generally have been in the range of 2-4 microns thick on articulating surfaces to avoid delamination and to minimize coating costs. Ceramic coatings applied by thin film technology to non-articulating surfaces are intended primarily to achieve biological compatibility, and hence have been at the lower end of this 2-4 micron range of coating thicknesses. Ceramic coatings have been applied to greater thickness by other coating technologies (e.g. plasma spray) primarily for other art areas, such as kitchen appliances. However, thick ceramic coatings applied by other technologies are also subject to cracking and delamination.
Despite the known hardness and wear resistance of thin film ceramic materials, the inventors herein have determined that thin film ceramic coatings applied to articulating surfaces tend to wear through well within the anticipated life of the patient and the prostheses. The wearing through of the prior art ceramic coating exposes the prior art metallic alloy substrate with the above described problems of wear debris, corrosion and biological incompatibility. The inventor's herein further believe that similar wear through occurs at other load bearing non-articulating surfaces due to the above described micromovement between the coated prosthesis and adjacent bone.
In view of the above, it is an object of the subject invention to provide an orthopedic prosthesis with enhanced wear resistance.
It is another object of the subject invention to provide an orthopedic prosthesis that ensures both wear resistance and biological compatibility.
A further object of the subject invention is to provide an orthopedic prosthesis that avoids galvanic corrosion between dissimilar metallic alloys of a prosthetic system.
Yet another object of the subject invention is to provide a coated orthopedic prosthesis that avoids cracking and delamination in use.